US7826587B1 - System and method of fast kVp switching for dual energy CT - Google Patents

System and method of fast kVp switching for dual energy CT Download PDF

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US7826587B1
US7826587B1 US12/558,248 US55824809A US7826587B1 US 7826587 B1 US7826587 B1 US 7826587B1 US 55824809 A US55824809 A US 55824809A US 7826587 B1 US7826587 B1 US 7826587B1
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kvp
snr
integration
ray source
period
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David Allen Langan
John Eric Tkaczyk
James Walter LeBlanc
Colin R. Wilson
Xiaoye Wu
Dan Xu
Thomas Matthew Benson
Jed Douglas Pack
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General Electric Co
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General Electric Co
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Priority to JP2010195242A priority patent/JP4726995B2/ja
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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/40—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment with arrangements for generating radiation specially adapted for radiation diagnosis
    • A61B6/405—Source units specially adapted to modify characteristics of the beam during the data acquisition process
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/02—Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/03—Computerised tomographs
    • A61B6/032—Transmission computed tomography [CT]
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/48—Diagnostic techniques
    • A61B6/482—Diagnostic techniques involving multiple energy imaging

Definitions

  • the present invention relates generally to diagnostic imaging and, more particularly, to an apparatus and method of acquiring imaging data at more than one energy range using a multi-energy imaging source.
  • an x-ray source emits a fan-shaped or cone-shaped beam toward a subject or object, such as a patient or a piece of luggage.
  • subject and object shall include anything capable of being imaged.
  • the beam after being attenuated by the subject, impinges upon an array of radiation detectors.
  • the intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the subject.
  • Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element.
  • the electrical signals are transmitted to a data processing system for analysis, which ultimately produces an image.
  • X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point.
  • X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom.
  • each scintillator of a scintillator array converts x-rays to light energy.
  • Each scintillator discharges light energy to a photodiode adjacent thereto.
  • Each photodiode detects the light energy and generates a corresponding electrical signal. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
  • a CT imaging system may include an energy sensitive (ES), multi-energy (ME), and/or dual-energy (DE) CT imaging system that may be referred to as an ESCT, MECT, and/or DECT imaging system, in order to acquire data for material decomposition or effective Z or monochromatic image estimation.
  • ESCT/MECT/DECT provides energy discrimination. For example, in the absence of object scatter, the system derives the material attenuation at a different energy based on the signal from two relative regions of photon energy from the spectrum: the low-energy and the high-energy portions of the incident x-ray spectrum. In a given energy region relevant to medical CT, two physical processes dominate the x-ray attenuation: (1) Compton scatter and the (2) photoelectric effect.
  • the detected signals from two energy regions provide sufficient information to resolve the energy dependence of the material being imaged. Furthermore, detected signals from the two energy regions provide sufficient information to determine the materials attenuation coefficients in terms of Compton scatter and photoelectric effect.
  • the material attenuation may be expressed as the relative composition of an object composed of two hypothetical materials, or the density and effective atomic number with the scanned object.
  • energy sensitive attenuation can be expressed in terms of two base materials, densities, effective Z number, or as two monochromatic representations having different keV.
  • Such systems may use a direct conversion detector material in lieu of a scintillator.
  • the ESCT, MECT, and/or DECT imaging system in an example is configured to be responsive to different x-ray spectra.
  • Energy sensitive detectors may be used such that each x-ray photon reaching the detector is recorded with its photon energy.
  • One technique to acquire projection data for material decomposition includes using energy sensitive detectors, such as a CZT or other direct conversion material having electronically pixelated structures or anodes attached thereto.
  • energy sensitive detectors such as a CZT or other direct conversion material having electronically pixelated structures or anodes attached thereto.
  • such systems typically include additional cost and complexity of operation in order separate and distinguish energy content of each received x-ray photon.
  • a conventional scintillator-based third-generation CT system may be used to provide energy separation measurements.
  • Such systems may acquire projections sequentially at different peak kilovoltage (kVp) operating levels of the x-ray tube, which changes the peak and spectrum of energy of the incident photons comprising the emitted x-ray beams.
  • kVp peak kilovoltage
  • a principle objective of scanning with two distinctive energy spectra is to obtain diagnostic CT images that enhance information (contrast separation, material specificity, etc.) within the image by utilizing two scans at different polychromatic energy states.
  • a number of techniques have been proposed to achieve energy sensitive scanning including acquiring two scans at, for instance, 80 kVp and 140 kVp (1) back-to-back sequentially in time where the scans require two rotations of the gantry around the subject that may be hundreds of milliseconds to seconds apart, (2) interleaved as a function of the rotation angle requiring one rotation around the subject, or (3) using a two tube/two detector system with the tubes/detectors mounted ⁇ 90 degrees apart, as examples.
  • taking separate scans several seconds apart from one another may result in mis-registration between datasets caused by patient motion (both external patient motion and internal organ motion) and different cone angles, and cannot be applied reliably where small details need to be resolved for body features that are in motion.
  • a ⁇ 90 degree separation in a two tube/two detector system inherently includes a mis-registration of datasets and adds cost and complexity to the overall system.
  • High frequency, low capacitance generators have made it possible to switch the kVp potential of the high frequency electromagnetic energy projection source on alternating views and interleave datasets.
  • data for two energy sensitive scans may be obtained in a temporally interleaved fashion rather than with separate scans made several seconds apart or with a two tube/two detector system.
  • such systems typically include a change to filament current to account for a changing mAs when kVp potential is switched.
  • the change in filament current can cause a change in filament temperature which, in turn, can cause a change in focal spot position and/or size.
  • Tube voltage may be used in establishing focal spot width with kVp switching, resulting in an oscillating focal spot width.
  • the change in focal spot position may be addressed through re-sampling of imaging data to mitigate the alignment issue. Or, if there is a significant change in focal spot size, a sinogram having the smaller focal spot may be blurred for improved registration between high and low kVp sinograms. However, these mitigation strategies tend to degrade resolution of the final image.
  • an x-ray source may be constructed having a pair of cathodes therein, each configured to emit electrons toward an anode, and each having a respective filament current associated therewith.
  • Such a system may accomplish fast kVp switching by, for instance, gridding the cathodes for the respective low and high kVp shots, with each cathode having a low and high kVp applied thereto relative to the anode. Though such a system may avoid the necessity of rapidly altering kVp or mA in a single cathode, it is at the expense of system complexity—both of hardware and system operation.
  • Embodiments of the invention are directed to a method and apparatus for acquiring imaging data at more than one energy range that overcome the aforementioned drawbacks.
  • a CT system includes a rotatable gantry having an opening for receiving an object to be scanned, an x-ray source coupled to the gantry and configured to project x-rays through the opening, a generator configured to energize the x-ray source to a first kVp and to a second kVp that is lower than the first kVp, a detector attached to the gantry and positioned to receive x-rays from the x-ray source that pass through the opening, and a controller configured to energize the x-ray source to the first kVp for a first time period subsequently energize the x-ray source to the second kVp for a second time period different from the first, integrate data from the detector for a first integration period that includes a portion of a steady-state period of the x-ray source at the first kVp, integrate data from the detector for a second integration period that includes a portion of a steady-state period of the x-ray source
  • a method of acquiring energy sensitive CT imaging data using a CT imaging system includes applying a first voltage potential to an x-ray source for a first time duration, applying a second voltage potential to the x-ray source for a second time duration that is greater than the first time duration, acquiring imaging data during a first integration period that includes when the x-ray source emits x-rays at a steady-state at the first potential, acquiring imaging data during a second integration period that includes when the x-ray source emits x-rays at a steady-state at the second potential, optimizing a first signal-to-noise ratio (SNR) during the first integration period with a second SNR during the second integration period by adjusting at least one operating parameter of the CT imaging system, and generating a dual-energy CT image using imaging data acquired after adjusting the at least one operating parameter of the CT imaging system.
  • SNR signal-to-noise ratio
  • a computer readable storage medium having stored thereon a computer program comprising instructions which when executed by a computer cause the computer to apply a first kVp potential to an x-ray source to obtain a first kVp steady-state operation of a CT imaging system, apply a second kVp potential to the x-ray source to obtain a second kVp steady-state operation of the CT imaging system, integrate a first set of imaging data that includes data obtained from x-rays generated during a time period when the x-ray source is at the first kVp steady-state operation, integrate a second set of imaging data that includes data obtained from x-rays generated during a time period when the x-ray source is at the second kVp stead-state operation, compare a first signal-to-noise ratio (SNR) of the integrated first set of imaging data with a second SNR of the integrated second set of imaging data, and adjust at least one operating parameter of the CT imaging system based on the comparison.
  • SNR signal-to-
  • a method of establishing first and second integration periods for acquisition of fast-switching dual-energy CT data in a CT system includes determining a first signal-to-noise ratio (SNR) for the first integration period, determining a second SNR for the second integration period, comparing the first SNR to the second SNR, and adjusting an operating condition of the CT system based on the compared first SNR and the second SNR.
  • SNR signal-to-noise ratio
  • FIG. 1 is a pictorial view of a CT imaging system.
  • FIG. 2 is a block schematic diagram of the system illustrated in FIG. 1 .
  • FIG. 3 is a perspective view of one embodiment of a CT system detector array.
  • FIG. 4 is a perspective view of one embodiment of a detector.
  • FIG. 5 is a flowchart illustrating acquisition of high and low kVp data sets, according to an embodiment of the invention.
  • FIG. 6 is a timing diagram illustrating high and low kVp data acquisition, according to an embodiment of the invention.
  • FIG. 7 is a pictorial view of a CT system for use with a non-invasive package inspection system according to an embodiment of the invention.
  • Diagnostics devices comprise x-ray systems, magnetic resonance (MR) systems, ultrasound systems, computed tomography (CT) systems, positron emission tomography (PET) systems, ultrasound, nuclear medicine, and other types of imaging systems.
  • Applications of x-ray sources comprise imaging, medical, security, and industrial inspection applications.
  • an implementation is employable for the detection and conversion of x-rays.
  • an implementation is employable for the detection and conversion of other high frequency electromagnetic energy.
  • An implementation is employable with a “third generation” CT scanner and/or other CT systems.
  • CT computed tomography
  • present invention is equally applicable for use with other multi-slice configurations, and with systems having a capability of shifting, or “wobbling” the focal spot during operation.
  • present invention will be described with respect to the detection and conversion of x-rays.
  • present invention is equally applicable for the detection and conversion of other high frequency electromagnetic energy.
  • present invention will be described with respect to a “third generation” CT scanner, but is equally applicable with other CT systems.
  • Embodiments of the invention support the acquisition of both anatomical detail as well as tissue characterization information for medical CT, and for components within luggage.
  • Energy discriminatory information or data may be used to reduce the effects of beam hardening and the like.
  • the system supports the acquisition of tissue discriminatory data and therefore provides diagnostic information that is indicative of disease or other pathologies.
  • This detector can also be used to detect, measure, and characterize materials that may be injected into the subject such as contrast agents and other specialized materials by the use of optimal energy weighting to boost the contrast of iodine and calcium (and other high atomic or materials).
  • Contrast agents can, for example, include iodine that is injected into the blood stream for better visualization.
  • the effective atomic number generated from energy sensitive CT principles allows reduction in image artifacts, such as beam hardening, as well as provides addition discriminatory information for false alarm reduction.
  • a computed tomography (CT) imaging system 10 is shown as including a gantry 12 representative of a “third generation” CT scanner.
  • Gantry 12 has an x-ray source 14 that projects a beam of x-rays 16 toward a detector assembly 18 that includes a collimator on the opposite side of the gantry 12 .
  • x-ray source 14 includes either a stationary target or a rotating target.
  • Detector assembly 18 is formed by a plurality of detectors 20 and data acquisition systems (DAS) 32 .
  • the plurality of detectors 20 sense the projected x-rays that pass through a medical patient 22 , and DAS 32 converts the data to digital signals for subsequent processing.
  • Each detector 20 produces an analog electrical signal that represents the intensity of an impinging x-ray beam and hence the attenuated beam as it passes through the patient 22 .
  • gantry 12 and the components mounted thereon rotate about a center of rotation 24 .
  • Control mechanism 26 includes an x-ray controller 28 and generator 29 that provides power and timing signals to x-ray source 14 and a gantry motor controller 30 that controls the rotational speed and position of gantry 12 .
  • An image reconstructor 34 receives sampled and digitized x-ray data from DAS 32 and performs high speed reconstruction. The reconstructed image is applied as an input to a computer 36 which stores the image in a mass storage device 38 .
  • Computer 36 also receives commands and scanning parameters from an operator via console 40 that has some form of operator interface, such as a keyboard, mouse, voice activated controller, or any other suitable input apparatus.
  • An associated display 42 allows the operator to observe the reconstructed image and other data from computer 36 .
  • the operator supplied commands and parameters are used by computer 36 to provide control signals and information to DAS 32 , x-ray controller 28 and gantry motor controller 30 .
  • computer 36 operates a table motor controller 44 which controls a motorized table 46 to position patient 22 and gantry 12 . Particularly, table 46 moves patients 22 through a gantry opening 48 of FIG. 1 in whole or in part.
  • System 10 may be operated in either monopolar or bipolar modes.
  • monopolar operation either the anode is grounded and a negative potential is applied to the cathode, or the cathode is grounded and a positive potential is applied to the anode.
  • bipolar operation an applied potential is split between the anode and the cathode.
  • monopolar or bipolar a potential is applied between the anode and cathode, and electrons emitting from the cathode are caused to accelerate, via the potential, toward the anode.
  • the cathode When, for instance, a ⁇ 140 kV voltage differential is maintained between the cathode and the anode and the tube is a bipolar design, the cathode may be maintained at, for instance, ⁇ 70 kV, and the anode may be maintained at +70 kV.
  • the cathode accordingly is maintained at this higher potential of ⁇ 140 kV while the anode is grounded and thus maintained at approximately 0 kV. Accordingly, the anode is operated having a net 140 kV difference with the cathode within the tube.
  • detector assembly 18 includes rails 17 having collimating blades or plates 19 placed therebetween. Plates 19 are positioned to collimate x-rays 16 before such beams impinge upon, for instance, detector 20 of FIG. 4 positioned on detector assembly 18 .
  • detector assembly 18 includes 57 detectors 20 , such as will be illustrated, each detector 20 having an array size of 64 ⁇ 16 of pixel elements 50 . As a result, detector assembly 18 has 64 rows and 912 columns (16 ⁇ 57 detectors) which allows 64 simultaneous slices of data to be collected with each rotation of gantry 12 .
  • detector 20 includes DAS 32 , with each detector 20 including a number of detector elements 50 arranged in pack 51 .
  • Detectors 20 include pins 52 positioned within pack 51 relative to detector elements 50 .
  • Pack 51 is positioned on a backlit diode array 53 having a plurality of diodes 59 .
  • Backlit diode array 53 is in turn positioned on multi-layer substrate 54 .
  • Spacers 55 are positioned on multi-layer substrate 54 .
  • Detector elements 50 are optically coupled to backlit diode array 53
  • backlit diode array 53 is in turn electrically coupled to multi-layer substrate 54 .
  • Flex circuits 56 are attached to face 57 of multi-layer substrate 54 and to DAS 32 .
  • Detectors 20 are positioned within detector assembly 18 by use of pins 52 .
  • x-rays impinging within detector elements 50 generate photons which traverse pack 51 , thereby generating an analog signal which is detected on a diode within backlit diode array 53 .
  • the analog signal generated is carried through multi-layer substrate 54 , through flex circuits 56 , to DAS 32 wherein the analog signal is converted to a digital signal.
  • FIG. 5 is a flowchart that illustrates a process 100 for obtaining optimized settings for a CT system to obtain fast kVp switching data, according to an embodiment of the invention.
  • desired steady-state high and low kVp settings are established based on a number of criteria.
  • the selection criteria are based on factors that include but are not limited to the object to be imaged, desired energy separation between high and low kVp, system capabilities, system rise and falltimes, and effective low and high kVps.
  • effective high and low kVp are to be distinguished from target and steady-state high kVp and low kVp.
  • the desired steady-state high kVp is 140 kVp
  • the desired low kVp is 80 kVp
  • other desired high and low kVps may be applied, depending on other factors as discussed.
  • the low kVp may, in that instance, be increased—but it may be at the expense of energy sensitivity or decreased energy separation between low and high kVp if the high kVp is not, likewise, increased.
  • an application is selected.
  • patient motion is inferred thereby specifying a minimum gantry rotation speed in order to “freeze” patient motion, as is understood with the art.
  • a neuroimaging scan is typically 1 second/revolution gantry speed
  • an abdominal scan is typically 0.5 second/revolution
  • a cardiac scan is typically at 0.35 second/revolution.
  • gantry rotation speeds may be selected based on any imaging application, based on system runspeed capabilities, patient size, and the like.
  • high and low kVp view times are established.
  • the view times established are based on a number of factors that include but are not limited to the gantry rotation speed as selected at step 104 and desired angular sampling rates (which is dependent on geometry of gantry 12 , geometry of detector assembly 18 , and the like).
  • High and low kVp view times are established with a constraint that their sum is a total view time that is established by gantry speed and system geometry.
  • gantry speed and system geometry establish a total view time, and each total view time includes a high and a low kVp viewtime, according to an embodiment of the invention.
  • each individual high and low kVp viewtime is each individual high and low kVp viewtime, but given the constraint that their sum comprises the total viewtime.
  • high and low integration times and trigger points are established at step 108 . High and low kVp viewtimes, integration times, and integration trigger points will be discussed and illustrated with respect to FIG. 6 .
  • timing diagram 200 illustrates a repeating pattern of high and low kVp operation according to an embodiment of the invention.
  • Timing diagram 200 includes a desired or target high kVp 202 and a desired or target low kVp 204 against a time axis 206 .
  • Timing diagram 200 illustrates high kVp shots 208 and low kVp shots 210 , two of which are illustrated, in a repeating pattern.
  • switching between high kVp shots 208 and low kVp shots 210 is not instantaneous, thus each transition from low kVp to high kVp includes a respective risetime 212 and a falltime 214 .
  • Risetime 212 and falltime 214 are on the order of 100 ⁇ s in one embodiment, but may be faster or slower than 100 ⁇ s depending on system hardware capacitance and other factors as understood within the art.
  • low kVp is illustrated beginning at steady-state at 216 .
  • target high kVp 202 is applied and a steady-state high kVp 220 is achieved, beginning at point 222 after risetime 212 .
  • target low kVp 204 is applied and a steady-state low kVp 226 is achieved, beginning at point 228 after falltime 214 .
  • Target high kVp 202 is again applied at point 230 , and the pattern of high kVp and low kVp operation repeats.
  • target high kVp 202 is applied during a high kVp duration 232 from point 218 , during risetime 212 and during steady-state high kVp 220 , to point 224 .
  • Low kVp is applied during a low kVp duration 234 from point 224 , during falltime 214 and during steady-state low kVp 226 , to point 230 , and the process repeats.
  • high and low kVp imaging data is integrated during the periods of high and low kVp operation.
  • high and low kVp triggering for integration occurs during respective rise and falltimes.
  • high kVp integration 236 begins at point 238 , which occurs during a portion of risetime 212 , during all of steady-state high kVp 220 operation, and during a portion of falltime 214 , to point 240 , when low kVp integration begins.
  • low kVp integration 242 begins at point 240 , which occurs during a portion of falltime 214 , during all of steady-state low kVp 226 operation, and during a portion of the subsequent risetime that begins at point 230 to point 244 .
  • trigger points 238 , 240 are selected based on a threshold voltage or on a selected time during respective rise and falltimes 212 , 214 .
  • integration for both high and low kVp acquisitions may be caused to occur during steady-state operation of high kVp, low kVp, or both.
  • no part of steady-state high kVp occurs during low kVp integration because, if steady state high kVp is integrated into a low kVp spectrum, material decomposition performance may be degraded.
  • an integrated average or effective high kVp 246 results that is somewhat lower than target high kVp 202 , because high kVp integration 236 occurs during portions of both risetime 212 and falltime 214 that are each lower in voltage than target high kVp 202 .
  • an integrated average or effective low kVp 248 results that is somewhat greater than target low kVp 204 , because low kVp integration 242 occurs during portions of falltime 214 and the subsequent risetime beginning at point 230 that are each greater in voltage than target low kVp 204 .
  • an effective energy difference 250 results between high kVp shots 208 and low kVp shots 210 , and effective energy difference 250 is dependent, for at least the reasons discussed, on the operating parameters of high and low kVp operation.
  • integration of rise and/or falltime may complicate integrated low and high kVp spectrums, which should be accounted for in a CT calibration procedure, during data correction, and during material decomposition processing.
  • step 106 includes establishing high and low kVp viewtimes, which in an example are at respective locations 232 and 234 of FIG. 6 .
  • view times for high and low kVp are selected that optimize acquired high and low kVp sinograms.
  • the ⁇ 1000 views each include a high kVp shot and a low kVp shot.
  • filament current is allowed to float during kVp switching.
  • mA realized is also a function of kVp.
  • commanded mA is constant, but despite this mA realized floats and therefore changes with applied kVp. Consequently, an approximate 1 ⁇ 3 drop in realized mA is typically experienced at low kVp compared to high kVp, despite a constant mA setting.
  • the low kVp generates generally lower energy x-rays, that are less penetrating. Accordingly, more mAs or mA times integration time is typically needed to achieve a desired signal.
  • the mA for the low kVp is maximized and therefore a longer integration time is required to achieve the desired mAs.
  • low kVp duration 234 is illustrated as significantly greater in duration than high kVp duration 232 .
  • kVp rise and falltimes are generally lower for high mA.
  • a relatively high mA is selected for high kVp operation and, due to mA dependence on kVp and tube filament temperature, mA for low kVp operation is approximately 2 ⁇ 3 of mA for high kVp operation.
  • viewtimes may be first selected based on the above parameters and conditions, and such initial settings may be based on the application (inferring patient motion), and a subject to be imaged.
  • Information regarding the subject to be imaged may be obtained a priori from a scout scan, from accumulated tables of imaging information, or from information obtained from the subject in a prior scan, as examples.
  • high and low kVp integration times and trigger points are established at step 108 of FIG. 5 , and may be based on anticipated signal levels or noise levels as understood in the art.
  • Resulting effective high and low kVps are determined at step 110 , which may be obtained in a manner consistent with timing diagram 200 of FIG. 6 .
  • the effective high and low kVps are influenced, as discussed, by risetimes and falltimes, by the portions of the rise and falltimes that are integrated into the view, and the target high and low steady-state kVps.
  • imaging parameters are selected in order to optimize a signal-to-noise ratio for a dual kVp imaging acquisition.
  • SNR signal-to-noise ratio
  • overall SNR may not be optimized for fast kVp switching based on parameters selected in steps 102 - 108 .
  • unnecessary and excess dose may be applied to a patient in order to obtain adequate SNR.
  • process 100 includes a number of iteration steps for optimizing SNR for high kVp (SNR H ) and SNR for low kVp (SNR L ) to avoid excess dose while realizing an adequate SNR for high and low kVp views.
  • An optimization function in material decomposition relates SNR H and SNR L to image quality of material density, and relates SNR H and SNR L to monochromatic representations, effective atomic number, and other image representations derived from high and low kVp acquisitions.
  • a goal is to balance SNR H and SNR L such that they are approximately equal to one another.
  • there are other optimization functions for optimizing SNR H and SNR L referring back to FIG. 5 , after effective high and low kVps are obtained at step 110 , an inquiry is made at step 112 as to whether SNR H and SNR L are optimized.
  • SNR H and SNR L are not optimized 114 , then one or a combination of system parameters are adjusted at step 116 , wherein the system parameters adjusted include, but are not limited to, gantry speed, high and low kVp view times, high and low kVp integration times, and high and low kVp trigger points.
  • system parameters adjusted include, but are not limited to, gantry speed, high and low kVp view times, high and low kVp integration times, and high and low kVp trigger points.
  • effective high and low kVps are again obtained at step 110 , and a determination is again made at step 112 as to whether SNR H and SNR L are optimized.
  • the iteration continues through steps 110 - 116 until SNR H and SNR L are optimized at step 112 , and, once optimized, high and low kVp imaging data is acquired having optimized SNR H and SNR L at step 118 .
  • step 104 in an iteration having a selected high and low kVp (step 102 ) and a selected gantry speed (step 104 ), wherein the iteration does not include changing gantry speed, because the gantry speed is unchanged, and because the number of views is unchanged, total integration time for high kVp integration 236 and low kVp integration 242 is unchanged.
  • trigger points for integration and/or integration times themselves may be iterated upon, however such is to be done with the constraint that total integration time is unchanged.
  • high kVp integration 236 time may be increased and low kVp integration 242 may correspondingly be decreased, however their sum, in this example, remains unchanged.
  • effective high and low kVps determined at step 110 may be altered, thus affecting SNR H with respect to SNR L in this example.
  • a controller such as controller 28 of FIG. 2 , is configured to generate an image using integrated data, according to an embodiment of the invention.
  • the controller is configured to generate the image using integrated data from multiple steady-state periods acquired during portions of the gantry rotation, each steady-state period containing one of the first kVp for the first time period and the second kVp for the second time period.
  • the controller is configured to calculate SNR H with SNR L and to adjust the operating parameter for all steady-state periods before a CT acquisition begins.
  • the controller is configured to adjust the operating parameter for multiple steady-state periods based in part on integration data in a previous steady-state period.
  • package/baggage inspection system 510 includes a rotatable gantry 512 having an opening 514 therein through which packages or pieces of baggage may pass.
  • the rotatable gantry 512 houses a high frequency electromagnetic energy source 516 as well as a detector assembly 518 having scintillator arrays comprised of scintillator cells similar to that shown in FIG. 4 or 5 .
  • a conveyor system 520 also is provided and includes a conveyor belt 522 supported by structure 524 to automatically and continuously pass packages or baggage pieces 526 through opening 514 to be scanned.
  • Objects 526 are fed through opening 514 by conveyor belt 522 , imaging data is then acquired, and the conveyor belt 522 removes the packages 526 from opening 514 in a controlled and continuous manner.
  • postal inspectors, baggage handlers, and other security personnel may non-invasively inspect the contents of packages 526 for explosives, knives, guns, contraband, etc.
  • An implementation of the system 10 and/or 510 in an example comprises a plurality of components such as one or more of electronic components, hardware components, and/or computer software components. A number of such components can be combined or divided in an implementation of the system 10 and/or 510 .
  • An exemplary component of an implementation of the system 10 and/or 510 employs and/or comprises a set and/or series of computer instructions written in or implemented with any of a number of programming languages, as will be appreciated by those skilled in the art.
  • An implementation of the system 10 and/or 510 in an example comprises any (e.g., horizontal, oblique, or vertical) orientation, with the description and figures herein illustrating an exemplary orientation of an implementation of the system 10 and/or 510 , for explanatory purposes.
  • An implementation of the system 10 and/or the system 510 in an example employs one or more computer readable signal bearing media.
  • a computer-readable signal-bearing medium in an example stores software, firmware and/or assembly language for performing one or more portions of one or more implementations.
  • An example of a computer-readable signal-bearing medium for an implementation of the system 10 and/or the system 510 comprises the recordable data storage medium of the image reconstructor 34 , and/or the mass storage device 38 of the computer 36 .
  • a computer-readable signal-bearing medium for an implementation of the system 10 and/or the system 510 in an example comprises one or more of a magnetic, electrical, optical, biological, and/or atomic data storage medium.
  • an implementation of the computer-readable signal-bearing medium comprises floppy disks, magnetic tapes, CD-ROMs, DVD-ROMs, hard disk drives, and/or electronic memory.
  • an implementation of the computer-readable signal-bearing medium comprises a modulated carrier signal transmitted over a network comprising or coupled with an implementation of the system 10 and/or the system 510 , for instance, one or more of a telephone network, a local area network (“LAN”), a wide area network (“WAN”), the Internet, and/or a wireless network
  • a CT system includes a rotatable gantry having an opening for receiving an object to be scanned, an x-ray source coupled to the gantry and configured to project x-rays through the opening, a generator configured to energize the x-ray source to a first kVp and to a second kVp that is lower than the first kVp, a detector attached to the gantry and positioned to receive x-rays from the x-ray source that pass through the opening, and a controller configured to energize the x-ray source to the first kVp for a first time period subsequently energize the x-ray source to the second kVp for a second time period different from the first, integrate data from the detector for a first integration period that includes a portion of a steady-state period of the x-ray source at the first kVp, integrate data from the detector for a second integration period that includes a portion of a steady-state period of the x-ray source
  • a method of acquiring energy sensitive CT imaging data using a CT imaging system includes applying a first voltage potential to an x-ray source for a first time duration, applying a second voltage potential to the x-ray source for a second time duration that is greater than the first time duration, acquiring imaging data during a first integration period that includes when the x-ray source emits x-rays at a steady-state at the first potential, acquiring imaging data during a second integration period that includes when the x-ray source emits x-rays at a steady-state at the second potential, optimizing a first signal-to-noise ratio (SNR) during the first integration period with a second SNR during the second integration period by adjusting at least one operating parameter of the CT imaging system, and generating a dual-energy CT image using imaging data acquired after adjusting the at least one operating parameter of the CT imaging system.
  • SNR signal-to-noise ratio
  • a computer readable storage medium having stored thereon a computer program comprising instructions which when executed by a computer cause the computer to apply a first kVp potential to an x-ray source to obtain a first kVp steady-state operation of a CT imaging system, apply a second kVp potential to the x-ray source to obtain a second kVp steady-state operation of the CT imaging system, integrate a first set of imaging data that includes data obtained from x-rays generated during a time period when the x-ray source is at the first kVp steady-state operation, integrate a second set of imaging data that includes data obtained from x-rays generated during a time period when the x-ray source is at the second kVp stead-state operation, compare a first signal-to-noise ratio (SNR) of the integrated first set of imaging data with a second SNR of the integrated second set of imaging data, and adjust at least one operating parameter of the CT imaging system based on the comparison.
  • SNR signal-to-
  • a method of establishing first and second integration periods for acquisition of fast-switching dual-energy CT data in a CT system includes determining a first signal-to-noise ratio (SNR) for the first integration period, determining a second SNR for the second integration period, comparing the first SNR to the second SNR, and adjusting an operating condition of the CT system based on the compared first SNR and the second SNR.
  • SNR signal-to-noise ratio
  • a technical contribution for the disclosed method and apparatus is that it provides for a computer-implemented apparatus and method of acquiring imaging data at more than one energy range using a multi-energy imaging source.
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